The present invention generally relates to methods and apparatus for generating an impedance cardiogram and, more particularly, to methods and apparatus for generating an impedance cardiogram for patients having cardiac stimulators which include bioimpedance sensors, as well as methods and apparatus for calibrating cardiac stimulators which include bioimpedance sensors.
For a variety of reasons, a person""s heart may not function properly and, thus, endanger the person""s well-being. Medical devices have been developed to assist physicians in the diagnosis and treatment of cardiac dysfunction. In regard to the diagnosis of cardiac dysfunction, it has been found that the volume of blood that a person""s heart is able to pump, commonly referred to as xe2x80x9ccardiac output,xe2x80x9d is one of the most important cardiovascular parameters. The cardiac output reflects the supply of oxygen and nutrients to the body. Measurements of cardiac output provide information for quantifying the extent of cardiac dysfunction and for indicating the optimal course of treatment.
Both invasive and non-invasive instruments are available for measuring a person""s cardiac output. The invasive techniques for measuring cardiac output require complex instrumentation, which must be operated by skilled personnel, and involve the penetration of the skin by a catheter. Due to the various disadvantages of invasive techniques, non-invasive techniques are generally preferred in the majority of cases.
Although a variety of non-invasive techniques exist, one of the more popular techniques is referred to as xe2x80x9cimpedance cardiography.xe2x80x9d Impedance cardiography is a method of measuring the electrical impedance of the body to determine cardiac output. In impedance cardiography, electrodes are typically connected at two locations on the body. A device, referred to as a cardiorespiratory monitor, generates an electric current that flows through the body from one electrode to the other. A second pair of electrodes, which are positioned between the first pair of electrodes, sense the potential developed by the electrical current as it flows through the body and delivers this sensed potential to the monitor. Based on the sensed potential and the injected current, the monitor calculates the impedance of the body.
In general, the impedance of the portion of the body between the electrodes varies inversely with the amount of blood flowing through the vessels in that region. Such impedance is often referred to as xe2x80x9cbioimpedancexe2x80x9d because it is the impedance of a set of biological tissues. In particular, if the first pair of electrodes are placed such that the current flows through the thorax, i.e., the cavity in which the heart and lungs lie, then the changes in the measured impedance result from changes in the amount of blood pumped by the heart.
The instantaneous amount of blood in the vessels is directly related to the performance of the heart. When blood is pumped out of the heart, the vessels in the thorax become momentarily filled with blood, and the impedance in the thorax rapidly decreases. After the ventricular contraction is complete, the impedance increases to its former level. Analysis of bioimpedance can therefore provide information related to cardiac output. Specifically, to obtain the cardiac output, the stroke volume, which is the amount of blood being ejected during each cardiac cycle, is first computed. The stroke volume may be calculated in a number of different ways, but, generally, it relates to the derivative of the impedance signal. Once the stroke volume has been determined, the cardiac output is computed by multiplying the stroke volume by the heart rate.
However, as anyone familiar with the bioelectrical characteristics of the human body is well aware, a variety of different factors can influence a bioimpedance measurement. For instance, one of the problems encountered in using thoracic impedance to derive the stroke volume is that the thoracic impedance is influenced by the effects of respiration. Similarly, if the patient is moving, during a stress test for instance, the movement also interferes with the thoracic impedance measurement and, thus, the subsequent calculation of stroke volume and cardiac output. Furthermore, when this technique is applied to patients suffering from severe cardiac dysfunction, the measured thoracic impedance may vary markedly from one cycle to another, thus making a qualitative determination of cardiac output difficult to obtain. In view of various problems such as these, a variety of different techniques have been developed for better correlating the measured thoracic impedance to cardiac output by eliminating the influences of these various problematic factors. As a result, impedance cardiography has improved vastly over the past several years and has become an important technique in the detection and treatment of cardiac dysfunction.
Bioimpedance signals are not only useful in the generation of impedance cardiographs using non-invasive monitors as discussed above. For instance, once a person has been diagnosed as having cardiac dysfunction, a physician may determine that a cardiac stimulator may be used to treat the condition. A cardiac stimulator is a medical device that delivers electrical stimulation to a patient""s heart. The cardiac stimulator generally includes a pulse generator for creating electrical stimulation pulses and a conductive lead for delivering these electrical stimulation pulses to the designated portion of the heaRT.
To understand how impedance measurement may be used to enhance the operation of a cardiac stimulator, it is beneficial to understand how cardiac stimulators have evolved. Early pacemakers did not monitor the condition of the heaRT. Rather, early pacemakers simply provided stimulation pulses at a fixed rate and, thus, kept the heart beating at that fixed rate. However, it was found that pacemakers of this type used an inordinate amount of energy because the stimulation pulses were not always needed. The human heart includes a sinus node located above the atria. The sinus node provides the electrical stimulation that causes a heart to contract. Even the sinus node of a heart in need of a pacemaker often provides such stimulation. Accordingly, if a heart, even for a short period, is able to beat on its own, providing an electrical stimulation pulse using a pacemaker wastes the pacemaker""s energy.
To conserve power, pacemakers were subsequently designed to monitor the heart and to provide stimulation pulses only when necessary. These pacemakers were referred to as xe2x80x9cdemandxe2x80x9d pacemakers because they provided stimulation only when the heart demanded stimulation. If a demand pacemaker detected a natural heartbeat within a prescribed period of time, typically referred to as the xe2x80x9cescape intervalxe2x80x9d, the pacemaker provided no stimulation pulse. Because monitoring uses much less power than generating stimulation pulses, the demand pacemakers took a large step toward conserving the limited energy contained in the pacemaker""s battery.
Clearly, the evolution of the pacemaker did not cease with the advent of monitoring capability. Indeed, the complexity of pacemakers has continued to increase in order to address the physiological needs of patients as well as the efficiency, longevity, and reliability of the pacemaker. For instance, even the early demand pacemakers provided stimulation pulses, when needed, at a fixed rate, such as 70 pulses per minute. To provide a more physiological response, pacemakers having a programably selectable rate were developed. So long as the heart was beating above this programably selected rate, the pacemaker did not provide any stimulation pulses. However, if the heart rate fell below this programably selected rate, the pacemaker sensed the condition and provided stimulation pulses as appropriate.
To provide even further physiological accuracy, pacemakers have now been developed that automatically change the rate at which the pacemaker provides stimulation pulses. These pacemakers are commonly referred to as xe2x80x9crate-responsivexe2x80x9d pacemakers. Rate-responsive pacemakers sense a physiological parameter of the patient and alter the rate at which the stimulation pulses are provided to the heart. Typically, this monitored physiological parameter relates to the changing physiological needs of the patient. For instance, when a person is at rest, the person""s heart may beat relatively slowly to accommodate the person""s physiological needs. Conversely, when a person is exercising, the person""s heart tends to beat rather quickly to accommodate the person""s heightened physiological needs. Unfortunately, the heart of a person in need of a pacemaker may not be able to beat faster on its own. In fact, prior to the development of rate-responsive pacemakers, patients were typically advised to avoid undue exercise, and pacemaker patients that engaged in exercise tended to tire quickly.
Rate-adaptive pacemakers help relieve this problem by sensing one or more physiological parameters of a patient that indicates whether the heart should be beating slower or faster. If the pacemaker determines that the heart should be beating faster, the pacemaker adjusts its base rate upward to provide a faster pacing rate if the patient""s heart is unable to beat faster on its own. Similarly, if the pacemaker determines that the patient""s heart should be beating more slowly, the pacemaker adjusts its base rate downward to conserve energy and to conform the patient""s heartbeat with the patient""s less active state.
One common rate-adaptive sensor measures physical activity as a parameter for rate adaption. Quite commonly, a cardiac stimulator employs an accelerometer, which is a device that responds to a patient""s movements, to measure the patient""s physical activity. One important advantage of a physical-activity sensor is its rapid response to patient activity. For instance, a physical-activity sensor responds favorably to activities which create vibration, such as jogging, walking, and stair climbing. Furthermore, the typical physical-activity sensor is quite simple in that it requires no special leads or implantation procedures.
A physical-activity sensor may possess various disadvantages as well. First, it should be understood that a physical-activity sensor is not generally regarded as a truly physiologic sensor because it does not measure true metabolic demand. For instance, activities such as bicycling may require an increased metabolic demand by the patient, but such activities may not promote rate adaptation because little vibration or few accelerations occur. Furthermore, cardiac stimulators that employ physical activity sensors typically begin rate adaption only when the vibration or accelerations measured by the physical activity sensor exceeds a preprogrammed level. Thus, it is difficult for such a cardiac stimulator to attain a scaled response to gradations of metabolic demand. Also, a physical activity sensor may generate undesirable responses to noise disturbances external to the body, such as vibrations caused by machinery, or from disturbances within the body, such as coughing, sneezing, and laughing.
As can be seen from the above-discussion, physical activity sensors, while exhibiting certain important advantages, do not provide the cardiac stimulator with a true measure of metabolic demand. Accordingly, certain rate-adaptive pacemakers have been developed which employ a metabolic-demand sensor that may be used alone or in conjunction with a physical activity sensor. A metabolic-demand sensor analyzes impedance signals that relate to cardiac performance to adapt the pacing rate to the metabolic demands of the patient. For instance, the pacemaker may analyze impedance measurements to determine the stroke volume of the heart and the minute volume of respiration as indications of the metabolic need of the patient. In normal human subjects with healthy hearts, the stroke volume of the heart has been found to remain relatively constant over a wide range of exertion. Increases in cardiac output required to meet physiologic needs are primarily provided by increasing the heart rate. However, when a patient with a pacemaker begins to exert himself, the heart attempts to increase its stroke volume to meet the increased metabolic needs of the patient. However, the stroke volume cannot increase by a factor of more than about 2 or 2xc2xd times for a typical patient. Thus, changes in stroke volume may be measured to increase the pacing rate to provide the increased cardiac output required by the increased metabolic demand of the patient. In addition to determining stroke volume and minute ventilation, cardiac stimulators may also use impedance sensors to derive similar rate and non-rate parameters from derived cardiac and hemodynamic parameters and, also, to detect and confirm tachyarrhythmias.
Much like the externally attached sensors described above in regard to the non-invasive cardiorespiratory monitor, impedance is measured in implantable devices by establishing an electrical current between two implanted electrodes and measuring the resulting voltage between the same electrodes or between another set of implanted electrodes. As with the external monitors, the current may be applied as a relatively high frequency AC carrier or as periodic narrow pulses. As a general rule, modern implantable cardiac stimulators use pulse-based methods rather than the high-frequency AC carrier methods because the former methods exhibit reduced power consumption in comparison with the latter methods.
Although metabolic-demand sensors of this type enhance the performance of rate adaptive pacemakers, certain difficulties still exist. For instance, the implantable devices use filters and other signal processing means to extract parameters from the impedance signals, and these parameters are believed to be representative of some physiological parameter. However, due to the wide variability between implants, empirical calibration factors are used to relate these estimated parameters to the true metabolic demand of the patient. Unfortunately, because of the interference described below, it is difficult to measure true metabolic demand with current cardiorespiratory monitors that use bioimpedance sensing.
Furthermore, the use of a non-invasive impedance cardiograph technique of a patient having an implanted rate adaptive cardiac stimulator which uses a metabolic demand sensor can create various problems. These problems generally originate because of the great similarity in impedance measurement techniques between the external impedance cardiograph device and the internal impedance-measuring cardiac stimulator. Because each device uses similar signals and similar portions of the elecromagnetic spectrum, simultaneous impedance measurement using both devices results in cross-interference. Such cross-interference may cause the implanted cardiac stimulator to operate in a less than ideal physiologic manner, and it may also cause the impedance cardiograph to generate inaccurate measurements of cardiac output.
The present invention may be applicable to one or more of the problems discussed above.
Certain aspects commensurate in scope with the originally claimed invention are set forth below. It should be understood that these aspects are presented merely to provide the reader with a brief summary of certain forms the invention might take and that these aspects are not intended to limit the scope of the invention. Indeed, the invention may encompass a variety of aspects that may not be set forth below.
In accordance with one aspect of the present invention, there is provided a cardiorespiratory measurement device that includes means for determining a property of a signal generated by an implanted device, and means for generating a bioimpedance sensing signal correlative to the property that produces substantially no interference with the signal of the implanted device.
In accordance with another aspect of the present invention, there is provided a cardiorespiratory measurement device that includes means for detecting signal properties of a pulsed signal emitted by a bioimpedance sensor of an implanted device, means for analyzing the detected signal properties to determine properties of an impedance sensing signal, and means for generating the impedance sensing signal that produces substantially no interference with the pulsed signal of the implanted device.
In accordance with still another aspect of the present invention, there is provided a cardiorespiratory monitor that includes a first pair of electrodes adapted to be coupled in spaced apart relation to one another on a patient""s body. A second pair of electrodes is adapted to be coupled in spaced apart relation to one another on the patient""s body. A voltage sensor is coupled to the first pair of electrodes to detect pulsed voltage signals within the patient""s body generated by a device implanted in the patient""s body. A processor is coupled to the voltage sensor. The processor receives the detected pulsed voltage signals and determines a frequency and timing of the detected pulsed voltage signals. The processor generates a timing signal correlative to the frequency and timing of the detected pulsed voltage signals. An impedance sensor control circuit is coupled to receive the timing signal from the processor. The impedance sensor circuit generates an impedance sensing signal correlative to the timing signal and delivers the impedance sensing signal to one of the second pair of electrodes. The impedance sensing signal is timed to produce substantially no interference with the pulsed voltage signals generated by the implanted device.
In accordance with yet another aspect of the present invention, there is provided a cardiorespiratory monitor that includes a detector adapted to detect a pulsed voltage signal generated by the implantable device. A processor is coupled to the detector. The processor receives the detected pulsed voltage signal and determines a frequency and timing of the detected pulsed voltage signal. The processor generates a timing signal correlative to the frequency and timing of the detected pulsed voltage signal. An impedance sensor control circuit is coupled to receive the timing original from the processor. The impedance sensor circuit generates an impedance sensing signal correlative to the timing signal. The impedance sensing signal is timed to produce substantially no interference with the pulsed voltage signal generated by the implanted device. A first pair of electrodes is adapted to be coupled in spaced apart relation to one another on a patient""s body. The first pair of electrodes delivers the impedance sensing signal to the patient""s body. A second pair of electrodes is adapted to be coupled in spaced apart relation to one another on the patient""s body. The second pair of electrodes detects the impedance sensing signal delivered by the first pair of electrodes and delivers the detected impedance sensing signal to the processor. The processor generates at least one signal correlative to the detected impedance sensing signal.